U.S. patent application number 12/380623 was filed with the patent office on 2009-08-27 for blood flow monitor with arterial and venous sensors.
This patent application is currently assigned to Thermal Technologies, Inc.. Invention is credited to H. Frederick Bowman.
Application Number | 20090216136 12/380623 |
Document ID | / |
Family ID | 46301063 |
Filed Date | 2009-08-27 |
United States Patent
Application |
20090216136 |
Kind Code |
A1 |
Bowman; H. Frederick |
August 27, 2009 |
Blood flow monitor with arterial and venous sensors
Abstract
A technique is disclosed for determining blood flow in a living
body by changing the thermal energy level in the venous blood flow
path and determining temperatures in both the venous and arterial
blood flow paths. Blood flow is calculated as a function of the
change in energy level and the temperature differences in the
venous and arterial blood flow paths.
Inventors: |
Bowman; H. Frederick;
(Needham, MA) |
Correspondence
Address: |
James L. Neal
185 N. West Temple, #207
Salt Lake City
UT
84103
US
|
Assignee: |
Thermal Technologies, Inc.
Cambridge
MA
|
Family ID: |
46301063 |
Appl. No.: |
12/380623 |
Filed: |
March 2, 2009 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
10809120 |
Mar 24, 2004 |
7527598 |
|
|
12380623 |
|
|
|
|
10364773 |
Feb 11, 2003 |
6913576 |
|
|
10809120 |
|
|
|
|
09733595 |
Dec 8, 2000 |
6565516 |
|
|
10364773 |
|
|
|
|
08946366 |
Oct 7, 1997 |
6203501 |
|
|
09733595 |
|
|
|
|
08106068 |
Aug 13, 1993 |
5797398 |
|
|
08946366 |
|
|
|
|
60458100 |
Mar 26, 2003 |
|
|
|
Current U.S.
Class: |
600/505 |
Current CPC
Class: |
A61B 5/029 20130101;
A61B 5/028 20130101 |
Class at
Publication: |
600/505 |
International
Class: |
A61B 5/02 20060101
A61B005/02 |
Claims
1. A system for quantifying blood flow in a living subject
comprising: (a) a thermal energy source for increasing blood
temperature at a location in the venous side of the circulatory
system during preselected time intervals; (b) a first sensor for
sensing blood temperature in the venous side of the circulatory
system where blood temperature is substantially unaffected by the
output of the thermal energy source; (c) a second sensor for
sensing blood temperature in the arterial blood flow path where the
blood temperature is affected by the output of the thermal energy
source; and (d) means responsive to the first and second sensors
and output of said thermal energy source for calculating a blood
flow related value as a function of outputs of the first and second
sensors when blood temperature in the arterial side of the
circulatory system is affected by output of said thermal energy
source and the outputs of the first and second sensors when blood
temperature in the arterial side of the circulatory system is not
substantially affected by said thermal energy source.
2. A system according to claim 1 further comprising a catheter
adapted to be introduced into the venous system for supporting one
or both of the first sensor and the thermal energy source.
3. A system according to claim 1 wherein said calculating means
comprises means for calculating blood flow as a function of the
output of the thermal energy source and the difference between: the
temperature difference sensed by the first and second sensors when
the blood temperature in the arterial side of the circulatory
system is affected by output of the thermal energy source and the
temperature difference sensed by the first and second sensors when
the thermal energy source is not activated.
4. A system according to claim 1 wherein the thermal energy source
comprises an energy source for increasing blood temperature in or
near the right atrium or vena cava and said calculating means
calculates a blood flow value that corresponds to cardiac
output.
5. A system for quantifying blood flow in the circulatory system of
a living subject comprising: (a) heating means adapted to be
located in the venous flow path for intermittently elevating the
temperature of blood in the arterial flow path; (b) means for
sensing blood temperature in the arterial flow path; (c) means for
providing a value corresponding to blood temperature at a location
in the venous flow path not affected by the output of said heating
means; and (d) means for calculating blood flow as a function of
the output of said heating means when elevating blood temperature,
the difference in the values from said sensing means and said
providing means when the temperature of blood is not elevated by
said heating means and the difference in the values from said
sensing means and said providing means when the temperature of
blood is elevated by said heating means.
6. A system according to claim 5 wherein said providing means
comprises a second means for sensing blood temperature.
7. A method for calculating blood flow in the blood flow path of a
living subject comprising: (a) changing a thermal energy
characteristic of blood at a site in the venous side of the
circulatory system in or near the right atrium or vena cava during
preselected time intervals; (b) detecting the temperature
difference between a location in the venous system substantially
unaffected by said changes introduced at the site and a selected
location in the arterial system where blood temperature is affected
by said changing step; (c) detecting the temperature difference
between the location in the venous system and the location in the
arterial system in the absence of said changing step; and (d)
calculating blood flow as a function of the temperature difference
of the first said detecting step, the temperature difference of the
second said detecting step and the change introduced by said
changing step.
Description
INTRODUCTION
[0001] This invention relates generally to techniques for measuring
blood flow in a body and, more particularly, to the use preferably
of one or more temperature sensors for measuring thermal energy
changes in the blood flowing through the heart and to the use of
unique data processing techniques in response thereto for
determining cardiac output.
BACKGROUND OF THE INVENTION
[0002] While the invention can be used generally to measure blood
flow at various locations in a body, it is particularly useful in
measuring blood flow in the heart so as to permit the measurement
of cardiac output. Many techniques for measuring cardiac output
have been suggested in the art. Exemplary thermodilution techniques
described in the technical and patent literature include: "A
Continuous Cardiac Output Computer Based On Thermodilution
Principles", Normann et al., Annals of Biomedical Engineering, Vol.
17, 1989; "Thermodilution Cardiac Output Determination With A
single Flow-Directed Catheter", Forrester, et al., American Heart
Journal, Vol. 83, No. 3, 1972; "Understanding Techniques for
Measuring Cardiac Output", Taylor, et al., Biomedical
Instrumentation & Technology, May/June 1990; U.S. Pat. No.
4,507,974 of M. L. Yelderman, issued Apr. 2, 1985; U.S. Pat. No.
4,785,823, of Eggers et al., issued on Nov. 22, 1988; and U.S. Pat.
No. 5,000,190, of John H. Petre, issued on Mar. 19, 1991.
[0003] A principal limitation in the quantification of cardiac
output is the existence of thermal fluctuations inherent in the
bloodstream. Previous methods work with those fluctuations while
observing the effects of an input signal to calculate cardiac
output. The invention described herein uses a differential
measurement technique to substantially eliminate the effect of the
thermal fluctuations, permitting the use of a minimal thermal input
signal, which allows frequent or continuous measurements.
[0004] It is desirable to obtain accurate cardiac output
measurements in an effectively continuous manner, i.e., several
times a minute, so that a diagnosis can be achieved more rapidly
and so that rapid changes in a patient's condition can be monitored
on a more continuous basis than is possible using current
techniques. Moreover, it is desirable to obtain instantaneous
measurements of the cardiac output on a beat-to-beat basis to
evaluate the relative changes which occur from beat to beat, as
well as to determine the presence of regurgitation.
BRIEF SUMMARY OF THE INVENTION
[0005] In accordance with general principal of the invention, blood
flow and/or cardiac output is determined rapidly, using a technique
by which an indicator substance, or agent, is introduced into the
bloodstream between a pair of detectors. The detectors are
sensitive to a parameter functionally related to the concentration
or magnitude in the bloodstream of the selected indicator agent.
The detectors are positioned apart by a distance functionally
sufficient to allow a measurement to be made of the differential
value of the selected parameter as it exists from time-to-time
between the two detectors. The indicator agent, for example, may be
a substance to change the pH of the blood, a fluid bolus carrying
thermal energy, or a substance to change a selected characteristic
of the blood, or the direct introduction of thermal energy, or the
like.
[0006] A determination is made of the difference in the values of
the selected blood parameter as it exists at the two detectors,
prior to the introduction of indicator agent (i.e., the first
differential value). The selected indicator agent is then
introduced in a predetermined magnitude. Then again a determination
is made of the difference in values of the selected parameter as it
exists at the locations of the two detectors (i.e., the second
differential value).
[0007] Blood flow or cardiac output, depending on the specific
location of the detectors, can then be determined as a function of
the difference between the first differential value and the second
differential value. Because the ultimate measurement of blood flow
or cardiac output is based on the difference of the differences,
the system operates effectively with the introduction of the
indicator agent in a very low magnitude. In turn, this allows
measurements to be made rapidly so that effectively continuous
measurements are obtained.
[0008] In accordance with a preferred embodiment of the invention,
for example, cardiac output can be determined rapidly and with low
levels of thermal energy input. To achieve such operation, in a
preferred embodiment, the technique of the invention uses a pair of
temperature sensors positioned at two selected locations within a
catheter which has been inserted into the path of the blood flowing
through the heart of a living body. The, sensors detect the
temperature difference between the two locations. Depending on the
location of the temperature sensors in the circulatory system, the
measured temperature difference varies over time. It has been
observed that when the temperature sensors are placed within the
heart, e.g., so that one sensor lies in the vena cava, for example,
and the second in the right ventricle or pulmonary artery, the
temperature difference varies in a synchronous manner with the
respiratory cycle.
[0009] Thus, in the preferred embodiment of the invention the
temperature difference over at least one respiratory cycle is
measured and averaged to provide an average temperature difference.
The averaging, or integrating, action effectively eliminates, as a
confounding factor in the determination of cardiac output, the
effect of instantaneous blood temperature fluctuations, such as
cyclical, respiratory-induced fluctuations.
[0010] To make such determinations, an average temperature
difference is first calculated over a time period of at least one
respiratory cycle in which no thermal energy is introduced into the
blood flow path. Thermal energy of a predetermined and relatively
low magnitude is then introduced into the blood flow path to
produce a heating action therein at a location between the two
temperature sensors. Once the temperature rise induced by the
heating stabilizes, the average temperature difference between the
two locations is again calculated from temperature difference
measurements over a time period of at least one respiratory cycle
at the higher temperature level. The difference between the average
temperature differences which occurs when the thermal energy is
turned on, referred to as the rising temperature change, is
determined. The difference between the average temperature
differences which occurs when the thermal energy is turned off,
referred to as the falling temperature change, is similarly
determined. The cardiac output is calculated as a function of the
thermal energy input and the rising and falling temperature
changes. Because a relatively low level of thermal energy is used
in making measurements, the overall sequence of determinations can
be safely repeated multiple times per minute, for example, so that
an effectively continuous, or quasi-continuous, determination of
cardiac output is obtained.
[0011] Further, a pair of sensors may be introduced into the
circulatory system at other locations in a subject. For example, a
catheter may be introduced into the venous side of the circulatory
system to locate an energy source (e.g.: a heater) in or near the
vena cava or within the heart. A sensor may be located in the
venous side of the circulatory system, upstream of the energy
source, anywhere that temperature is not materially affected by
output from the energy source. This sensor is one means for
providing a reference signal that compensates for fluctuations in
bloodstream temperature introduced by factors other than the energy
source. That is, the reference signal compensates for background
bloodstream temperature. The other sensor is introduced into the
arterial side of the circulatory system anywhere it senses blood
temperature as affected by the energy source. It may be introduced
by a second catheter, for example an arterial catheter. Blood flow
and/or cardiac output may then be calculated as described
above.
[0012] In accordance with a further embodiment of the invention, a
temperature sensor that also acts as a source of thermal energy,
e.g., a thermistor, is positioned at a third location in the
cardiac blood flow path. Power is supplied to the sensor sufficient
to elevate the temperature of the sensor from a first temperature
level to a second temperature level. In one embodiment of the
invention, the temperature of the sensor is changed from the first
to the second level and is maintained constant at said second level
by varying the power that is supplied thereto. Such varying power
is proportional to the instantaneous flow velocity and, hence,
assuming a constant flow area, is proportional to the instantaneous
cardiac output. Measurement of the sensor heating power and the
temperature increment at the sensor can thus be used to
continuously effect a determination of the instantaneous cardiac
output. Further, for example, when the sensor is placed downstream
at the outlet of one of the heart chambers, the variation in flow
output over the cardiac cycle can be analyzed to provide an
indication of the regurgitation characteristics of the heart outlet
valve over the cardiac cycle. Moreover, such instantaneous cardiac
output determination can be further refined to compensate for
fluctuations in the temperature of the blood flowing through the
heart by measuring the instantaneous temperature of the blood with
another temperature sensor at a nearby location and appropriately
taking into account such temperature variations when determining
the cardiac output.
[0013] In another application, both the continuous cardiac output
determinations and the instantaneous cardiac output determinations,
as described above, can be combined. Thus, three temperature
sensors and a source of thermal energy can all be used in
combination to simultaneously provide an accurate and effectively
continuous determination of time-averaged cardiac output, and a
determination of instantaneous cardiac output at each instant of
the cardiac cycle in still another application, two temperature
sensors and a source of thermal energy can be used in an
appropriate sequence to provide the averaged cardiac output
determination and the instantaneous cardiac output
determination.
BRIEF DESCRIPTION OF THE DRAWINGS
[0014] The invention is described in detail with the help of the
accompanying drawings wherein:
[0015] FIG. 1 shows a simplified diagrammatic view of a human
heart;
[0016] FIG. 2 shows a simplified diagrammatic view of a catheter
useful in the invention;
[0017] FIG. 3 shows a flow chart depicting steps in a process used
in the invention;
[0018] FIG. 3A shows a smooth temperature difference curve obtained
in the process depicted in FIG. 3;
[0019] FIG. 4 shows a flow chart depicting further steps in a
process used in the invention;
[0020] FIG. 5 shows a flow chart depicting still further steps in a
process of the invention;
[0021] FIG. 6 shows a graph depicting a temperature/time relation
used in the invention;
[0022] FIG. 7A shows a simplified diagrammatic view of another
catheter used in the invention;
[0023] FIG. 7B shows a flow chart depicting steps in still another
process of the invention;
[0024] FIGS. 8A, 8B and 8C show flow charts depicting modes of
operating the process of the invention;
[0025] FIGS. 9A, 9B, and 9C show graphs of parameter relationships
used in the invention; and
[0026] FIG. 10 shows a graph useful for calibrating the flow values
for a catheter used in the invention.
[0027] FIG. 11 is a diagram of another catheter configuration.
DESCRIPTION OF THE INVENTION
[0028] As can be seen in FIG. 1, which represents a human heart 10
in a much simplified diagrammatic form, a flexible catheter 11 is
inserted through the veins into the right atrium, or auricle, 12 of
the heart and, thence, through the right ventricle 13 until the end
of the catheter resides in or near the exit, or pulmonary, artery
14 which leads to the lungs. As is well known, blood flows (as
represented by the arrows) from the input vein 15, i.e., the vena
cava, into the right atrium and right ventricle and thence
outwardly to the lungs and subsequently returns from the lungs into
the left atrium 16, through the left ventricle 17 and thence
outwardly into the aorta 18.
[0029] In accordance with the embodiment of the invention, shown
with reference to FIG. 1, temperature sensors, e.g., thermistors,
are carried by the catheter so that, when inserted as shown in FIG.
1, a first sensor 19 is positioned at a location within the vena
cava 15 or right atrium 12 and a second sensor 20 is positioned at
a location in or near the pulmonary artery 14.
[0030] For simplicity, the flexible catheter 11 is depicted in FIG.
2 in an extended condition with temperature sensors 19 and 20 at
two different locations for measuring temperatures T.sub.1 and
T.sub.2, respectively. A power source 21 of thermal energy which is
borne, or carried, by the catheter 11 is positioned in the right
atrium at a location between sensors 19 and 20. In a particular
embodiment, the catheter-borne source is, for example, a coil of
resistive wire placed on or embedded in the surface of catheter 11,
to which an AC or a DC voltage (not shown) at a controllable level
is supplied so as to generate thermal energy, i.e. heat. The
magnitude of the thermal energy can be suitably controlled to
insert a predetermined amount of thermal energy at a selected time,
which thermal energy is transferred to the blood flowing through
the heart so as to raise its temperature. The energy source is
positioned at a sufficient distance from the sensor 19 that the
latter is effectively thermally isolated from the site of the
thermal energy source.
[0031] While the locations of the sensors 19 and 20 and the energy
source 21 can be as shown in FIG. 1, alternative locations can also
be used. Thus, the sensor 19 can be positioned in the vena cava 15,
while the energy source 21 is located in the right atrium 12 and
the sensor 20 in either the right atrium or the right ventricle.
Moreover, if sensor 19 is positioned in the vena cava 15, the
entire energy source 21, which is normally elongated, need not be
located in the right atrium and can have a portion thereof in the
vena cava and a portion thereof in the right atrium. Such source
should preferably be at least partially located in the right
atrium. Further, sensor 20 may be positioned in the right ventricle
near the pulmonary artery 14 or may be located in the pulmonary
artery itself at or near the right ventricle.
[0032] The temperatures T.sub.1 and T.sub.2 at locations 19 and 20
upstream and downstream, respectively, from the thermal energy
source 21 are monitored and processed appropriately by a digital
microprocessor. In accordance with the invention, the instantaneous
temperatures are obtained as the outputs T.sub.1(t) and T.sub.2(t)
of the temperature sensors 19 and 20, respectively. The outputs are
connected to a differential amplifier to generate an analog signal
which is proportional to the temperature difference
.DELTA.T(t)=T.sub.1(t)-T.sub.2(t) between them. The temperature
difference signal .DELTA.T(t) is digitized and sampled at selected
time intervals by an analog-to-digital/sampling circuit. The
digitized sampled temperature difference values and the known
thermal energy values are supplied to a digital microprocessor
which then suitably processes the data to provide the desired
cardiac output information. The processing stages used in the host
microprocessor are implemented by suitable programming of the
microprocessor and are discussed below with the help of FIGS.
3-6.
[0033] The source 21 of thermal energy is alternately turned on and
off. If it is assumed that thermal stability is reached after each
change and that there is a substantially constant rate of blood
flow, a stable temperature difference can be measured in each case.
The quantity of blood flowing past the thermal energy source, i.e.,
the cardiac output, can be derived from such temperature difference
measurements. However, such derivation is complicated by two
factors which may affect the measurement of blood flow. First, the
rate of blood flow through the heart is not substantially constant
but surges with each heart contraction. Second, the temperature of
the blood flowing through the heart is not constant but varies with
each respiratory (breathing) cycle. In a preferred embodiment, the
processing of the data takes such factors into account, as
discussed below.
[0034] The process for determining cardiac output is performed in a
microprocessor 2 which in a first embodiment is programmed to
respond to the temperatures sensed at T.sub.1 and T.sub.2 and to
perform the steps depicted in accordance with the flow charts shown
in FIGS. 3-5. From a knowledge of such flow charts, it would be
well within the skill of those in the art to appropriately program
any suitable and known digital microprocessor, such as a personal
computer, to perform the steps shown.
[0035] FIG. 3 depicts a basic process, identified as Process I,
which is used in the overall processing of temperature data for
determining cardiac output, as subsequently depicted in FIGS. 4 and
5. In the basic process shown in FIG. 3, a temperature difference
as a function of time .DELTA.T(t) is determined by a differential
amplifier which responds to T.sub.1(t) and T.sub.2(t). Such
differences may be effectively smoothed, or filtered, to produce a
smooth temperature difference curve, as shown in FIG. 3A, which
varies as a function of time in a cyclic manner which depends
principally on the respiratory cycle of the person whose cardiac
output is being determined.
[0036] The periods .tau..sub.1, .tau..sub.2 . . . .tau..sub.n for
each respiratory cycle are determined over n cycles. A
characteristic of the temperature difference at each cycle is
determined. For example, such characteristic preferably is the
averaged temperature difference during each cycle
(.DELTA.T.tau..sub.1, .DELTA.T.tau..sub.2 . . . .DELTA.T.tau..sub.2
. . . .DELTA.T.tau..sub.n).
[0037] (Alternatively, for example, the peak temperature
differences may be the determined characteristic.) These averaged
temperature differences (.DELTA.T.sub..tau.n) are added for the n
cycles involved and are divided by n to determine an averaged
temperature difference per cycle (.DELTA.T.sub..tau.n). The use of
Process I is depicted in the process steps shown in FIG. 4,
identified as Process II.
[0038] As seen therein, the steps of Process I are first performed
when the source 21 of thermal energy (i.e., a heater) is turned off
and the average .DELTA.T.sub.off value per cycle is determined and
suitably stored. The sampling time at which such determination is
made is depicted in FIG. 6 as the sample time period S1.
[0039] The heater 21 is then turned on for a specific time period
to supply a known amount of power P to the blood flowing through
the heart and, accordingly, the temperature of the blood flowing
past the heater rises and the temperature difference .DELTA.T(t)
rises over a transition, or delay, rise time period, t.sub.R1,
shown in FIG. 6 and designated as D1, after which the temperature
difference generally stabilizes over a second sample time period
S2. As seen in FIG. 4, after the heater 21 is turned on and the
temperature has stabilized, Process I is performed, again over n
cycles, e.g., over the time period S2, and the averaged temperature
difference .DELTA.T.sub.on is determined with the heater turned on
and is suitably stored. The heater is then turned off and the
temperature falls over a transition, or delay, fall time period
t.sub.F1, shown in FIG. 6 and designated as D2, generally to its
former value.
[0040] Cardiac output is calculated using the averaged temperature
differences when the energy is off and the averaged temperature
differences when the energy is on, by the relationship:
F = P C p ( .DELTA. T _ on - .DELTA. T _ off ) ##EQU00001##
where: [0041] F=Flow [0042] P=Power [0043] C.sub.P=heat capacitance
[0044] .DELTA.T.sub.on=average temperature for power on
[0045] .DELTA.T.sub.off=average temperature for power off.
[0046] As seen in FIG. 5, the steps of Process II are repeated
indefinitely for N data collection cycles, a data collection cycle
being designated as including the time periods S1, D1, S2, and D2,
as shown in FIG. 6. For each data collection cycle the rise time
temperature difference .DELTA.TR between the averaged temperature
difference .DELTA.T.sub.on at S2 and the averaged temperature
difference .DELTA.T.sub.off at S1 and the fall time temperature
difference .DELTA.T.sub.F between the averaged temperature
difference .DELTA.T.sub.off at S1 and the averaged temperature
difference .DELTA.T.sub.on at S2 are determined.
[0047] The flow, F.sub.R, is calculated for each data collection
cycle from the known amount of power P introduced into the blood
flow stream by the energy source, or heater 21, from the known heat
capacitance of blood, C.sub.P, and from the difference in the
averaged temperature differences .DELTA.T.sub.on and
.DELTA.T.sub.off, which occurs over the data collection cycle
S1+D1+S2 in accordance with the following relationship:
F R = P C p ( .DELTA. T _ on - .DELTA. T _ off ) ##EQU00002##
[0048] In a similar manner, the flow F.sub.F is calculated from P,
C.sub.P and the difference in the averaged temperature differences
.DELTA.T.sub.on and .DELTA.T.sub.off which occurs over the later
portion of the data collection cycle S1+D1+S2 in accordance with
the following relationship:
F F = P C p ( .DELTA. T _ on - .DELTA. T _ off ) ##EQU00003##
[0049] F.sub.R and F.sub.F can be averaged to obtain the averaged
flow ( F) over one data collection cycle as shown in FIG. 6.
F _ = F R + F F 2 ##EQU00004##
[0050] A suitable calibration constant can be used to adjust the
values of F.sub.R, F.sub.F and F.
[0051] Accordingly, by using two temperature sensors 19 and 20,
cardiac output can be determined several times a minute in
accordance with FIGS. 3-6, yielding an effectively continuous
cardiac output value. Because such measurements can be made using
relatively low power levels, the danger that the heart may be
damaged by the introduction of thermal energy is substantially
eliminated.
[0052] It will be apparent that the foregoing technique, which has
been described in connection with the direct introduction of heat
as an indicator agent and the measurement of temperature, can
readily be performed by those skilled in the art by using indicator
agents which affect the pH of the blood or change other blood
parameters.
[0053] The system can be located in the body of a subject as
illustrated in the diagrammatic representation of the circulatory
system shown in FIG. 11. The circulatory system illustrated there
includes systemic venous system 101, vena cava 150, heart 100,
systemic arterial system 102, pulmonary venous system 106 and the
body's capillary system 216. The representation of the capillary
system 216 is intended to indicate that system anywhere in the
body. The representation of the heart 100 includes the right atrium
(RA), the right ventricle (RV), the left atrium (LA) and the left
ventricle (LV).
[0054] An energy source 210 and a sensor 190 are located in the
venous side 101 of the circulatory system of the subject by any
suitable placement method or means. For example, a catheter 110 may
be introduced into the vena cava 150 to locate the energy source
210 in or near the vena cava or within the heart 100 and the sensor
190 somewhere in the venous system upstream of the energy source
210.
[0055] FIG. 11 shows the energy source 210 in the vena cava 150
while FIG. 1 shows an energy source 21 in the right atrium 12. The
energy sources 21 and 210 are similar in function and may be
similar in construction. Sensor 190 is located upstream of the
energy source 210 where it is not materially affected by output
from the energy source, where it senses blood temperature
unaffected by the energy source. Sensor 190 or other suitable
reference provides a reference signal that compensates for
temperature level and fluctuations in bloodstream introduced by
factors other than the energy source 210. That is, the reference
signal corresponds to background bloodstream temperature,
unperturbed by the energy source 210. FIG. 11 shows one preferred
location for sensor 190 in or near the vena cava and an alternate
location more remote from the energy source. In another embodiment
of the system a fixed resistor or other reference value that
approximates the unperturbed blood temperature could be used, in
lieu of the value provided by the sensor 190, to compensate for
background bloodstream temperature. (This likely would reduce the
accuracy of the compensation and its use would depend on the
requirements of the user.) Sensor 190 can be similar in
construction and is similar in function to sensor 19 shown in FIG.
1.
[0056] The sensor 200 is located in the arterial system 102 and may
be introduced by any suitable method or means. For example, an
arterial catheter may be used. The sensor 200 may be located
anywhere on the arterial side 102 of the body where it senses blood
temperature affected by the energy source 210. The sensor 200 can
be located, for example, in the arm 104, leg or neck. The sensor
200 and the sensor 20 shown in FIG. 1 can be similar in
construction and are similar in function.
[0057] When the system is configured as shown in FIG. 11, cardiac
output and blood flow may be calculated as previously described.
When the sensor 200 and the energy source 210 are separated by such
a distance that a material amount of thermal energy from the energy
source is lost to the body before reaching sensor 200, a
compensating factor is included in the calibration constant
referenced in FIG. 5.
[0058] In some situations it may be desirable to provide more
frequent indications of cardiac output, such as, for example, the
instantaneous cardiac output or the cardiac output averaged over
each individual cardiac cycle (i.e. each heart beat). Such
information can be provided using the further embodiments of the
invention discussed below with reference to FIGS. 7-8. A single
temperature sensor 30 at a location near the distal end of the
catheter 31 (as shown in FIG. 7A) can be used to determine the
instantaneous or beat-to-beat blood velocity V(t). The blood
velocity can be combined with the cardiac output averaged over one
or more data collection cycles to calculate instantaneous cardiac
output. The process used is shown in the process depicted in FIG.
7B, identified as Process IV.
[0059] As seen therein, the initial temperature T.sub.3i(t) sensed
at temperature sensor 30 as a function of time is smoothed, or
filtered, in the manner as previously discussed above, and suitably
measured and stored at an initial time t.sub.o. A predetermined
rise in temperature .DELTA.T.sub.3 of the temperature sensor itself
is selected. Power is then supplied at time t.sub.o to the
temperature sensor 30 from a power source 30A connected thereto to
cause its temperature T.sub.3(t) to rise by a predetermined
amount.
[0060] Power may be supplied to the sensor in different ways
according to the needs of the particular measurement and the
relative simplicity or complexity of the required circuitry, three
such ways being depicted in FIGS. 8A, 8B, and 8C.
[0061] For example, in a first mode of operation (FIG. 8A), heating
power may be supplied to the sensor in such a manner as to keep the
final sensor temperature T.sub.3f(t) constant at an initial level
.DELTA.T.sub.3 above the initial temperature T.sub.3i(t.sub.o),
i.e. T.sub.3f=T.sub.3i(t.sub.o)+.DELTA.T.sub.3 even when the local
blood temperature varies with time, as illustrated in FIG. 9A.
Under such conditions, the sensor is maintained at a time-varying
temperature increment .DELTA.T.sub.3(t) above the instantaneous
local blood temperature, T.sub.b(t).
[0062] Alternatively, in a second mode of operation, power can be
supplied to the sensor so as to continuously maintain the sensor at
a fixed temperature increment above the time varying local blood
temperature, as illustrated in FIG. 9B. Under such conditions,
.DELTA.T.sub.3(t)=.DELTA.T.sub.3, a constant, and the sensor
temperature varies according to
T.sub.3f(t)=.DELTA.T.sub.3+.DELTA.T.sub.3i(t).
[0063] A third mode of heating may also be convenient when the
temperature sensors are temperature-sensitive resistors, or
thermistors. Thus, when a thermistor is used, it may be more
convenient to design an electrical heating circuit that maintains
the sensor at a constant resistance increment above the resistance
of the sensor that corresponds to the local blood temperature. If R
is the corresponding resistance for a sensor temperature T, then
these conditions are represented by
.DELTA.R.sub.3(t)=.DELTA.R.sub.3, a constant, and the sensor
resistance varies according to
R.sub.3f(t)=.DELTA.R.sub.3+.DELTA.R.sub.3i(t), as illustrated in
FIG. 9C. The change in temperature .DELTA.T.sub.3(t) is then
replaced by the change in resistance R.sub.3(t) in the ratio which
is integrated over a cardiac cycle. Further details and exemplary
apparatus for such modes of operation are presented and described
in U.S. Pat. No. 4,059,982, issued to E. F. Bowman on Nov. 29,
1977. With all three of the above approaches, power (P) is supplied
to produce a temperature rise (.DELTA.T) both of which are then
related to the instantaneous blood velocity and, hence, blood
flow.
[0064] Techniques in which sensor heating power and temperature can
be measured and used to provide more detailed information on
cardiac output are described below. The technique involved can be
applied to measure both instantaneous cardiac output, and the
cardiac output for an individual cardiac cycle. Such detailed
measurement information greatly enhances the diagnostic capability
of a physician.
[0065] First, a method is described to measure instantaneous
volumetric flow (which flow if measured at the location described
above is the cardiac output). For each of the particular
implementations described above, the power P.sub.3(t) applied to
the temperature sensor 30 is controlled so as to maintain the final
temperature of the sensor at a desired value T.sub.3f. The power
applied to the temperature sensor 30 or, more generally, the ratio
of the power applied to the sensor to the temperature increment,
P.sub.3(t)/.DELTA.T.sub.3(t), is directly correlated with the fluid
and flow properties of the flowing liquid about the sensor.
[0066] For example, the relationship between required sensor power
and local fluid velocity, V(t), is given by a correlation of the
form:
P(t)=4.pi.ka.DELTA.T(t)[1+C.sub.1P.sub.r.sup.n(2a.rho.V(t)/.mu.).sup.m]
[0067] Where [0068] P(t)=instantaneous power to sensor [0069]
k=thermal conductivity of fluid [0070] a=sensor radius [0071]
.DELTA.T(t)=instantaneous temperature difference between heated
sensor and unheated fluid temperature. [0072] C.sub.1=constant of
calibration [0073] P.sub.r=a nondimensional "Prandtl" number which
relates to the viscosity .mu., heat capacity [0074] Cp and thermal
conductivity k of a fluid. [0075] n,m=power factors which are
determined from experimental data [0076] .rho.=fluid density [0077]
.mu.=viscosity [0078] V(t)=instantaneous fluid velocity
[0079] The fluid flow velocity in the vicinity of the sensor can be
determined from the required sensor heating power. Volumetric flow
in the vessel can then be determined with one further assumption
for the distribution of the fluid flow within the vessel. For
example, assuming a uniform velocity profile within the vessel,
volumetric flow F.sub.3 is given by
F.sub.3=VA
where V is the fluid velocity in the vessel and A is the flow
area.
[0080] If the fluid flow area A is not previously known, it may be
inferred from the measurement of average volumetric flow in the
vessel. Such average volumetric flow can be determined, for
example, by using the techniques of the invention already described
above herein or by using other techniques for yielding comparable
information. For example, if F is the average cardiac output,
typically measured over several cardiac cycles, as described above,
and V is the average fluid velocity, determined by calculating an
average value for the instantaneous flow velocity over at least one
cardiac cycle, then one such estimate for the average flow area A
is given by
= F/ V
[0081] Therefore, given the sensor measured heating power, first
the fluid velocity and then volumetric flow can be calculated at
any desired instant in time, i.e., F.sub.3(t)=V(E) , yielding an
instantaneous measure of volumetric flow, i.e., cardiac output.
[0082] In another embodiment, a method to measure cardiac output
over a single cardiac cycle is described. As described above in
different implementations, the power P.sub.3(t) applied to the
temperature sensor 30 is controlled so as to maintain the
temperature of the sensor at a desired signal value T.sub.3f. The
power applied to the temperature sensor 30, or more generally, as
discussed above, the ratio of the power applied to the sensor to
the temperature increment, i.e., P.sub.3(t)/.DELTA.T.sub.3(t), is
directly correlated with the properties of the fluid flow in the
vicinity of the sensor.
[0083] Thus, the integrated value of the power to temperature ratio
over a single cardiac cycle is directly correlated with, i.e., is
proportional to the average cardiac output over the cardiac
cycle,
.intg. cardiac cycle P 3 ( t ) .DELTA. T 3 ( t ) t .varies. F _
##EQU00005##
or, alternatively expressed
1 T .intg. cardiac cycle P 3 ( t ) .DELTA. T 3 ( t ) t .varies. F _
##EQU00006##
where T represents the period of the cardiac cycle and F
correspondingly represents cardiac output averaged over the cardiac
cycle. Thus, the average cardiac output F over an individual
cardiac cycle then can be determined from the measured and
integrated power and temperature signals from the sensor.
[0084] Furthermore, an explicit correlation for integrated power
and average cardiac output over the cardiac cycle may be dispensed
with if a simple qualitative indication of the change in cardiac
output on a cardiac cycle-to-cardiac cycle basis is desired. To
obtain such information, a given measurement of cardiac output is
taken as associated with a corresponding measured value of the
integrated sensor signal over a cardiac cycle. The measurement of
cardiac output could be obtained intermittently by the techniques
described in this invention or other similar techniques. Since
cardiac output is known to be correlated with the value of the
integrated sensor signal over the cardiac cycle, any changes in the
sensor signal indicate a corresponding change in cardiac
output.
[0085] In certain situations, it may be desirable to compensate for
temperature variations in the blood which is flowing past the
sensor, as this may affect the value of F(t). A process for such
compensation is depicted as Process IV in FIG. 7B wherein a
temperature T.sub.2(t) is sensed by a second sensor (which may be,
for example, sensor 19 or sensor 20) at a location remote from
sensor 30 (see FIG. 7A). For example, knowledge of the
instantaneous blood temperature is required for the process in
which the heated sensor is maintained at a constant increment above
the local blood temperature. In this case, the temperature
T.sub.2(t) is used as a proxy for the temperature T.sub.3(t) which
would be measured in the absence of sensor heating.
[0086] In a further alternative embodiment, where only two sensors
19 and 20 are utilized (as shown in FIG. 2), sensor 20 can be used
as the primary sensor when calculating instantaneous cardiac output
(equivalent to sensor 30 in FIG. 7A) and sensor 19 can be used as
the secondary temperature compensation sensor. In such an
embodiment, the averaged cardiac output can be determined using
sensors 19 and 20, as set forth in FIGS. 3-6 and the instantaneous
cardiac output can then subsequently be determined using sensors 19
and 20, as set forth in FIG. 7B and either FIG. 8A, 8B or 8C, such
average and instantaneous cardiac output determinations being made
in sequence by the microprocessor to provide the cardiac
information in both forms, as desired.
[0087] As mentioned above, when using the above described catheter,
the various flow values which are determined in accordance with the
processes as discussed above are proportional to flow but may not
be equal to the actual flow values unless they are suitably
calibrated since the correspondence between the calculated and
actual values depends on the manner in which a particular catheter
is constructed and used. A calibration constant for a particular
catheter can be represented by the slope and intercept of a curve
which relates the calculated flow and the actual flow, in
accordance with the following relationship:
F.sub.actual=aF.sub.calc.+b
where, as illustrated in FIG. 10, "a" is the slope of a straight
line 35 and "b" is the intercept thereof along the vertical axis.
Curve 35 can be obtained by using a known catheter and known flow
values therein to construct a curve 36. The best straight line fit
is determined as line 35. The slope "a" and intercept "b" are
thereby determined. Such determined values for "a" and "b" can be
used with the calculated flow values in each case to determine the
actual flow from the calculated flow. While the above description
discusses preferred embodiments of the invention, modifications
thereof may occur to those in the art within the spirit and scope
of the invention. Hence, the invention is not to be construed as
limited to particular embodiments described, except as defined by
the appended claims.
* * * * *